Fiber reinforced composite stents

ABSTRACT

Polymeric composite stents reinforced with fibers for implantation into a bodily lumen are disclosed.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a continuation of application Ser. No. 12/878,937,filed on 9 Sep. 2010, now U.S. Pat. No. 8,741,201, which is a divisionof application Ser. No. 11/205,254, filed on 15 Aug. 2005, nowabandoned, all of which are incorporated by reference as if fully setforth, including any figures, herein.

BACKGROUND OF THE INVENTION

1. Field of the Invention

This invention relates to radially expandable implantable medicaldevices such as stents for implantation into a bodily lumen. Inparticular, the invention relates composite stents reinforced withfibers.

2. Description of the State of the Art

This invention relates to radially expandable endoprostheses, which areadapted to be implanted in a bodily lumen. An “endoprosthesis”corresponds to an artificial device that is placed inside the body. A“lumen” refers to a cavity of a tubular organ such as a blood vessel.

A stent is an example of such an endoprosthesis. Stents are generallycylindrically shaped devices, which function to hold open and sometimesexpand a segment of a blood vessel or other anatomical lumen such asurinary tracts and bile ducts. Stents are often used in the treatment ofatherosclerotic stenosis in blood vessels. “Stenosis” refers to anarrowing or constriction of the diameter of a bodily passage ororifice. In such treatments, stents reinforce body vessels and preventrestenosis following angioplasty in the vascular system. “Restenosis”refers to the reoccurrence of stenosis in a blood vessel or heart valveafter it has been treated (as by balloon angioplasty, stenting, orvalvuloplasty) with apparent success.

The treatment of a diseased site or lesion with a stent involves bothdelivery and deployment of the stent. “Delivery” refers to introducingand transporting the stent through a bodily lumen to a region, such as alesion, in a vessel that requires treatment. “Deployment” corresponds tothe expanding of the stent within the lumen at the treatment region.Delivery and deployment of a stent are accomplished by positioning thestent about one end of a catheter, inserting the end of the catheterthrough the skin into a bodily lumen, advancing the catheter in thebodily lumen to a desired treatment location, expanding the stent at thetreatment location, and removing the catheter from the lumen.

In the case of a balloon expandable stent, the stent is mounted about aballoon disposed on the catheter. Mounting the stent typically involvescompressing or crimping the stent onto the balloon. The stent is thenexpanded by inflating the balloon. The balloon may then be deflated andthe catheter withdrawn. In the case of a self-expanding stent, the stentmay be secured to the catheter via a retractable sheath or a sock. Whenthe stent is in a desired bodily location, the sheath may be withdrawnwhich allows the stent to self-expand.

The stent must be able to satisfy a number of mechanical requirements.First, the stent must be capable of withstanding the structural loads,namely radial compressive forces, imposed on the stent as it supportsthe walls of a vessel. Therefore, a stent must possess adequate radialstrength. Radial strength, which is the ability of a stent to resistradial compressive forces, is due to strength and rigidity around acircumferential direction of the stent. Radial strength and rigidity,therefore, may also be described as, hoop or circumferential strengthand rigidity.

Once expanded, the stent must adequately maintain its size and shapethroughout its service life despite the various forces that may come tobear on it, including the cyclic loading induced by the beating heart.For example, a radially directed force may tend to cause a stent torecoil inward. Generally, it is desirable to minimize recoil.

In addition, the stent must possess sufficient flexibility to allow forcrimping, expansion, and cyclic loading. Longitudinal flexibility isimportant to allow the stent to be maneuvered through a tortuousvascular path and to enable it to conform to a deployment site that maynot be linear or may be subject to flexure. Finally, the stent must bebiocompatible so as not to trigger any adverse vascular responses.

The structure of a stent is typically composed of scaffolding thatincludes a pattern or network of interconnecting structural elementsoften referred to in the art as struts or bar arms. The scaffolding canbe formed from wires, tubes, or sheets of material rolled into acylindrical shape. Conventional methods of constructing a stent from apolymer material involve extrusion, blow molding, or injection molding apolymer tube based on a single polymer or polymer blend and then lasercutting a pattern into the tube. The scaffolding is designed so that thestent can be radially compressed (to allow crimping) and radiallyexpanded (to allow deployment). A conventional stent is allowed toexpand and contract through movement of individual structural elementsof a pattern with respect to each other.

Additionally, a medicated stent may be fabricated by coating the surfaceof either a metallic or polymeric scaffolding with a polymeric carrierthat includes an active or bioactive agent or drug. Polymericscaffolding may also serve as a carrier of an active agent or drug.

Furthermore, it may be desirable for a stent to be biodegradable. Inmany treatment applications, the presence of a stent in a body may benecessary for a limited period of time until its intended function of,for example, maintaining vascular patency and/or drug delivery isaccomplished. Therefore, stents fabricated from biodegradable,bioabsorbable, and/or bioerodable materials such as bioabsorbablepolymers should be configured to completely erode only after theclinical need for them has ended.

In general, there are several important aspects in the mechanicalbehavior of polymers that affect stent design. Polymers tend to havelower strength than metals on a per unit mass basis. Therefore,polymeric stents typically have less circumferential strength and radialrigidity than metallic stents of the same or similar dimensions.Inadequate radial strength potentially contributes to a relatively highincidence of recoil of polymeric stents after implantation into vessels.

Another potential problem with polymeric stents is that their struts orbar arms can crack during crimping and expansion, especially for brittlepolymers. The localized portions of the stent pattern subjected tosubstantial deformation tend to be the most vulnerable to failure.Furthermore, in order to have adequate mechanical strength, polymericstents may require significantly thicker struts than a metallic stent,which results in an undesirably larger profile.

Additionally, another factor to consider in stent design is radiopacity.In addition to meeting the mechanical requirements described above, itis desirable for a stent to be radiopaque, or fluoroscopically visibleunder x-rays. “Radiopaque” refers to the ability of a substance toabsorb x-rays. Accurate stent placement is facilitated by real timevisualization of the delivery of a stent. A cardiologist orinterventional radiologist can track the delivery catheter through thepatient's vasculature and precisely place the stent at the site of alesion. This is typically accomplished by fluoroscopy or similar x-rayvisualization procedures. For a stent to be fluoroscopically visible itmust be more absorptive of x-rays than the surrounding tissue.Radiopaque materials in a stent may allow for its direct visualization.

A significant shortcoming of polymers as compared to metals (andpolymers generally composed of carbon, hydrogen, oxygen, and nitrogen)is that they are radiolucent with no radiopacity. Polymers tend to havex-ray absorption similar to body tissue.

Additionally, there are manufacturing difficulties in placing smallmarkers on stents as well as challenges in keeping very small markersattached to the stent. If the maximum permissible size of the marker istoo small to be visible on a fluoroscope, multiple markers may benecessary. This makes manufacturing even more challenging.

Therefore, it would be desirable to have methods of making biodegradablepolymeric stents that are both strong and flexible.

SUMMARY OF THE INVENTION

Certain embodiments of the present invention are directed to a method ofmaking a stent that may include forming a mixture having a matrixpolymer and a plurality of short fibers such that fibers include amaterial having a melting temperature greater than a melting temperatureof the matrix polymer. The method may further include disposing themixture in a tube or sheet forming apparatus to form a tube or a sheetsuch that the apparatus is heated so that a temperature of the mixturein the apparatus is greater than the melting temperature of the matrixpolymer and less than the melting temperature of the material of thefibers. At least a portion of the matrix polymer may be a polymer melt.A stent may be fabricated from the tube or sheet including the matrixpolymer and the short fibers.

Further embodiments of the present invention are directed to a method ofmaking a stent that may include forming a tube having at least one fiberlayer and at least one polymer film layer such that fibers of at leastone fiber layer include a material having a melting temperature greaterthan a melting temperature of at least one polymer film layer. Themethod may further include heating the tube to a temperature greaterthan the melting temperature of at least one polymer film layer and lessthan the melting temperature of the material of the fibers to melt atleast a portion of the polymer of at least one polymer film layer. Atleast a portion of at least one fiber layer may become embedded withinat least a portion of the melted polymer of at least one polymer filmlayer. The heated tube may then be cooled and a stent fabricated fromthe cooled tube.

Additional embodiments of the present invention are directed to a methodof making a stent that may include forming a layered sheet having atleast one fiber layer and at least one polymer film layer such thatfibers of at least one fiber layer include a material having a meltingtemperature greater than a melting temperature of at least one polymerfilm layer. The method may further include heating the layered sheet toa temperature greater than the melting temperature of at least onepolymer film layer and less than the melting temperature of the materialof the fibers to melt at least a portion of the polymer of at least onepolymer film layer. At least a portion of the fibers may become embeddedwithin at least a portion of the melted polymer of at least one polymerfilm layer. The heated layered sheet may then be cooled and a stentfabricated from the cooled sheet.

Additional embodiments of the present invention are directed to a methodof making a stent that may include forming a coating layer comprising acoating polymer over a tube-shaped fiber layer having a plurality offibers. The coating layer may be formed by applying a fluid includingthe coating polymer dissolved in a solvent and by removing all or amajority of the solvent from the applied fluid. The fibers may include amaterial that is insoluble or having a relatively low solubility in thesolvent. The material may have a melting temperature greater than amelting temperature of the coating polymer. The method may furtherinclude fabricating a stent from the coated fiber layer.

Other embodiments of the present invention are directed to a method ofmaking a stent that may include disposing a plurality of fibers within amold for forming a structure. The method may further include disposing amatrix polymer that is partially or completely molten into the mold toat least partially embed the fibers within the molten polymer. The fibermay include a material having a melting temperature greater than amelting temperature of the matrix polymer. A temperature of the matrixpolymer and the fibers in the mold may be less than the meltingtemperature of the material. The method may further include cooling themolten polymer to form the structure and fabricating a stent from thecooled structure.

Further embodiments of the present invention are directed to a method ofmaking a stent that may include disposing a plurality of fibers in anextruder for forming a structure. The method may further includeconveying a matrix polymer into the extruder. The fibers may include amaterial having a melting temperature greater than a melting temperatureof the matrix polymer. The structure may be formed with the extruder ata temperature greater than the melting temperature of the matrix polymerand less than the melting temperature of the material in such a way thatat least some of the fibers become embedded within the matrix polymer. Astent may then be fabricated from the cooled structure.

Additional embodiments of the present invention are directed to a methodof making a stent that may include heating a fiber mesh tube includingtwo types of fibers. A first fiber may include a first polymer and thesecond fiber may include a second polymer. The first polymer may have asoftening temperature lower than a softening temperature of the secondpolymer. The tube may be heated to a temperature range between thesoftening temperature of the first polymer and the softening temperatureof the second polymer. The method may further include applying pressureto the tube so as to flatten at least some of the fibers of the tube toreduce a radial profile of the tube.

Some further embodiments of the present invention are directed to amethod of making a stent that may include heating a fiber mesh tube. Atleast some of the fibers of the tube may include a first polymer and asecond polymer. The first polymer may have a softening temperature lowerthan a softening temperature of the second polymer. The tube may beheated to a temperature range between the softening temperature of thefirst polymer and the softening temperature of the second polymer. Themethod may further include applying pressure to the tube so as toflatten at least some of the fibers of the tube to reduce a radialprofile of the tube.

Some further embodiments of the present invention are directed to amethod of making a stent that may include coupling a metallic film to atleast a portion of a surface of a polymeric tube. The method may furtherinclude fabricating a stent from the tube with the metallic film so thatthe metallic film is over at least a portion of a surface of the stent.

Other embodiments of the present invention are directed to a method ofmaking a stent that may include forming a tube having a metallic film inbetween two radial polymeric layers. The method may further includefabricating a stent from the tube.

Certain other embodiments of the present invention are directed to amethod of making a stent that may include elongating a polymeric tube sothat a diameter of the stent decreases. The method may further includepositioning a metallic band around a circumference of the elongatedtube. The elongated polymeric tube with the metallic band positionedaround the tube may then be heated. The method may further includeallowing the heated tube to radially expand so as to couple the metallicband to the tube. A stent may be fabricated from the expanded tube.

Additional embodiments of the present invention are directed to aradially expandable stent including a plurality of interconnectingstructural elements including fibers at least partially embedded in amatrix polymer. The fibers may include a material having a meltingtemperature greater than a melting temperature of the matrix polymer.The fibers may be configured to provide mechanical reinforcement to thestent due to a higher strength and modulus along an axis of the fibersthan the matrix polymer.

Other embodiments of the present invention are directed to a radiallyexpandable stent including a plurality of interconnecting structuralelements including at least one radial fiber layer and at least oneradial polymer film layer. The fibers may include material with amelting temperature greater than a melting temperature than at least onepolymer film layer. At least one fiber layer may be at least partiallyembedded within at least one polymer film layer. The fibers may beconfigured to provide mechanical reinforcement to the stent due to ahigher strength and modulus along an axis of the fibers than the polymerfilm layer.

Additional embodiments of the present invention are directed to aradially expandable stent including a plurality of structural elementsincluding at least two radial fiber layers and at least one radialpolymer film layer. At least a portion of at least one fiber layer maybe embedded within at least a portion of at least one polymer filmlayer. An orientation of fibers relative to a cylindrical axis of thestent of at least one fiber layer may be different from an orientationof fibers in another fiber layer.

Certain embodiments of the present invention are directed to a radiallyexpandable stent woven from at lease two types of fibers. A first fibermay include a first polymer and the second fiber may include a secondpolymer. The first polymer may have a softening temperature lower than asoftening temperature of the second polymer. At least some of the fibersmay have a flattened radial profile that reduces the radial profile ofthe tube.

Other embodiments of the present invention are directed to a radiallyexpandable stent woven from fibers comprising a first polymer and asecond polymer. The first polymer may have a softening temperature lowerthan a softening temperature of the second polymer such that at leastsome of the fibers have a flattened radial profile that reduces theradial profile of the tube.

Additional embodiments of the present invention are directed to aradially expandable stent including a metallic film coupled to aplurality of portions of a surface of the stent such that the metallicfilm is sufficiently radiopaque to allow the stent to be visualizedduring use

Further embodiments of the present invention are directed to a radiallyexpandable stent including a plurality of interconnecting structuralelements such that the structural elements may have two radial polymericlayers with metallic film embedded in a plurality of locations inbetween the layers. The metallic film may be sufficiently radiopaque toallow the stent to be visualized during use.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 depicts a three-dimensional view of a stent.

FIG. 2 depicts a schematic plot of the rate of crystallization of apolymer as a function of temperature.

FIG. 3 depicts a schematic representation of a mixture of a continuouspolymer phase and a discrete fiber phase.

FIG. 4 depicts a fiber-reinforced tube with short fibers.

FIG. 5 depicts an embodiment of a method of fabricating a fiberreinforced tube.

FIG. 6A depicts a two-dimensional radial cut-off view of a tube formedwith a fiber layer and two polymer layers.

FIG. 6B depicts an expanded view of the layers from FIG. 6A.

FIG. 7 depicts a tube of helically wound fiber mesh.

FIG. 8 depicts a two-dimensional view of layers of a tube formed withfiber layers and polymer layers.

FIG. 9 depicts a fiber mesh tube disposed on a mandrel.

FIG. 10 depicts alignment of struts or structural elements with fibers.

FIG. 11 depicts a radial cross-section of a composite fiber.

FIG. 12 depicts a radial cross-section of a system for heating andflattening fibers of a fiber stent.

FIG. 13 depicts an expanded view of fibers in the system of FIG. 12prior to flattening the fibers.

FIG. 14 depicts an expanded view of fibers in the system of FIG. 12showing flattening of the fibers.

FIG. 15 depicts a polymeric tube with a circumferentially alignedmetallic band coupled or adhered to the surface of the tube.

FIG. 16 depicts a polymeric tube with a longitudinally aligned strip ofmetallic film coupled or adhered to the surface of the tube.

FIG. 17 depicts a stent with a circumferentially aligned metallic filmon its surface.

FIG. 18 depicts a stent with a longitudinally aligned metallic film onits surface.

FIG. 19 depicts a cross-sectional view of a sidewall of a portion of astructural element of a stent with a coating above a metallic film on apolymeric substrate.

FIG. 20 depicts a cross-sectional view of a sidewall of a portion of astructural element of a stent with a metallic film embedded between anabluminal layer and a luminal layer.

DETAILED DESCRIPTION OF THE INVENTION

Various embodiments of the present invention relate to compositepolymeric biodegradable implantable medical devices and methods ofmaking such devices. In general, a composite implantable medical deviceis a device which is made up of two or more macroscopically distinctmaterials that have different properties. The composite device as awhole may have desirable properties of two or more of the distinctmaterials. Therefore, desirable mechanical and/or degradation propertiesmay be obtained through the use of a polymer composite structure.

For the purposes of the present invention, the following terms anddefinitions apply:

The “glass transition temperature,” T_(g), is the temperature at whichthe amorphous domains of a polymer change from a brittle vitreous stateto a solid deformable or ductile state at atmospheric pressure. In otherwords, the T_(g) corresponds to the temperature where the onset ofsegmental motion in the chains of the polymer occurs. When an amorphousor semicrystalline polymer is exposed to an increasing temperature, thecoefficient of expansion and the heat capacity of the polymer bothincrease as the temperature is raised, indicating increased molecularmotion. As the temperature is raised the actual molecular volume in thesample remains constant, and so a higher coefficient of expansion pointsto an increase in free volume associated with the system and thereforeincreased freedom for the molecules to move. The increasing heatcapacity corresponds to an increase in heat dissipation throughmovement. T_(g) of a given polymer can be dependent on the heating rateand can be influenced by the thermal history of the polymer.Furthermore, the chemical structure of the polymer heavily influencesthe glass transition by affecting mobility.

“Stress” refers to force per unit area, as in the force acting through asmall area within a plane. Stress can be divided into components, normaland parallel to the plane, called normal stress and shear stress,respectively. Tensile stress, for example, is a normal component ofstress applied that leads to expansion (increase in length). Inaddition, compressive stress is a normal component of stress applied tomaterials resulting in their compaction (decrease in length). Stress mayresult in deformation of a material, which refers to change in length.“Expansion” or “compression” may be defined as the increase or decreasein length of a sample of material when the sample is subjected tostress.

“Strain” refers to the amount of expansion or compression that occurs ina material at a given stress or load. Strain may be expressed as afraction or percentage of the original length, i.e., the change inlength divided by the original length. Strain, therefore, is positivefor expansion and negative for compression.

“Solvent” is defined as a substance capable of dissolving or dispersingone or more other substances or capable of at least partially dissolvingor dispersing the substance(s) to form a uniformly dispersed mixture atthe molecular- or ionic-size level. The solvent should be capable ofdissolving at least 0.1 mg of the polymer in 1 ml of the solvent, andmore narrowly 0.5 mg in 1 ml at ambient temperature and ambientpressure. The “strength” of a solvent refers to the degree to which asolvent may dissolve a polymer. The stronger a solvent is, the morepolymer the solvent can dissolve.

Furthermore, a property of a material that quantifies a degree of strainwith applied stress is the modulus. “Modulus” may be defined as theratio of a component of stress or force per unit area applied to amaterial divided by the strain along an axis of applied force thatresults from the applied force. For example, a material has both atensile and a compressive modulus. A material with a relatively highmodulus tends to be stiff or rigid. Conversely, a material with arelatively low modulus tends to be flexible. The modulus of a materialdepends on the molecular composition and structure, temperature of thematerial, amount of deformation, and the strain rate or rate ofdeformation. For example, below its T_(g), a polymer tends to be brittlewith a high modulus. As the temperature of a polymer is increased frombelow to above its T_(g), its modulus decreases.

“Above” a surface or layer is defined as higher than or over a surfaceor layer measured along an axis normal to a surface or layer, but notnecessarily in contact with the surface or layer.

“Vicat Softening Temperature” (VST) is a measure of the temperature atwhich a polymer starts to soften at specified test conditions accordingto ISO 306. It is determined with a standard indenter (a flat-endedneedle of 1 mm 2 circular cross section) penetrating into the surface ofa test specimen under a predefined load. The temperature at 1 mmpenetration is quoted as the VST in Co. VST gives an indication of amaterial's ability to withstand limited short-term contact with a heatedobject.

The term “elastic deformation” refers to deformation of an object inwhich the applied stress is small enough so that the object movestowards its original dimensions or essentially its original dimensionsonce the stress is released. However, an elastically deformed polymermaterial may be prevented from returning to an undeformed state if thematerial is below the T_(g) of the polymer. Below T_(g), energy barriersmay inhibit or prevent molecular movement that allows deformation orbulk relaxation.

“Elastic limit” refers to the maximum stress that a material willwithstand without permanent deformation. The “yield point” is the stressat the elastic limit and the “ultimate strain” is the strain at theelastic limit. The term “plastic deformation” refers to permanentdeformation that occurs in a material under stress after elastic limitshave been exceeded.

The term “implantable medical device” is intended to include, but notlimited to, self-expandable stents, balloon-expandable stents,stent-grafts, and grafts. In general, an implantable medical device,such as a stent, can have virtually any structural pattern that iscompatible with a bodily lumen in which it is implanted. The embodimentsof the invention described herein are generally applicable toimplantable medical devices.

Typically, a stent is composed of a pattern or network ofcircumferential rings and longitudinally extending interconnectingstructural elements of struts or bar arms. In general, the struts arearranged in patterns, which are designed to contact the lumen walls of avessel and to maintain vascular patency. A myriad of strut patterns areknown in the art for achieving particular design goals.

FIG. 1 depicts a three-dimensional view of a stent 10 which shows struts15. The embodiments disclosed herein are not limited to stents or to thestent pattern illustrated in FIG. 1. The embodiments are easilyapplicable to other patterns and other devices. The variations in thestructure of patterns are virtually unlimited.

A stent such as stent 10 may be fabricated from a tube by forming apattern with a technique such as laser cutting. Representative examplesof lasers that may be used include an excimer, carbon dioxide, and YAG.In other embodiments, chemical etching may be used to form a pattern onthe elongated tube.

In some embodiments, the diameter of the polymer tube prior tofabrication of an implantable medical device may be between about 0.2 mmand about 5.0 mm, or more narrowly between about 1 mm and about 3 mm.Unless otherwise specified, the “diameter” of the tube refers to theoutside diameter of the tube.

Various embodiments of fabricating composite polymeric implantabledevices are disclosed herein. The continuous and discrete phases mayinclude polymeric or metallic materials or a combination of polymericand metallic materials.

In general, polymers can be biostable, bioabsorbable, biodegradable, orbioerodable. Biostable refers to polymers that are not biodegradable.The terms biodegradable, bioabsorbable, and bioerodable, as well asdegraded, eroded, and absorbed, are used interchangeably and refer topolymers that are capable of being completely eroded or absorbed whenexposed to bodily fluids such as blood and can be gradually resorbed,absorbed, and/or eliminated by the body. In addition, a medicated stentmay be fabricated by coating the surface of the stent with an activeagent or drug, or a polymeric carrier including an active agent or drug.

A stent made from a biodegradable polymer is intended to remain in thebody for a duration of time until its intended function of, for example,maintaining vascular patency and/or drug delivery is accomplished. Afterthe process of degradation, erosion, absorption, and/or resorption hasbeen completed, no portion of the biodegradable stent, or abiodegradable portion of the stent will remain. In some embodiments,very negligible traces or residue may be left behind. The duration canbe from about a month to a few years, but is typically in the range ofsix to eighteen months.

Biodegradation of polymers generally refers to changes in physical andchemical properties that occur in a polymer upon exposure to bodilyfluids as in a vascular environment. The changes in properties mayinclude a decrease in molecular weight, deterioration of mechanicalproperties, and decrease in mass due to erosion or absorption.Mechanical properties may correspond to strength and modulus of thepolymer. Deterioration of the mechanical properties of the polymerdecreases the ability of a stent, for example, to provide mechanicalsupport in a vessel. The decrease in molecular weight may be caused by,for example, hydrolysis and/or metabolic processes. Hydrolysis is achemical process in which a molecule is cleaved into two parts by theaddition of a molecule of water.

Consequently, the degree of bulk degradation of a polymer is stronglydependent on the diffusivity, and hence the diffusion rate of water inthe polymer. Several characteristics or parameters of the degradationprocess are important in designing biodegradable devices. These includean average erosion rate of a device, the erosion profile, the half-lifeof the degrading polymer, and mechanical stability of a device duringthe degradation process. The “average erosion rate” may be an averageerosion rate over any selected time interval:Average erosion rate=(m ₂ −m ₁)/(t ₂ −t ₁)where “m” refers to mass of the device, “t” refers to a time duringerosion, and m₁ and m₂ are the masses of the device at t₁ and t₂ duringerosion. For instance, the selected time interval may be between theonset of degradation and another selected time. Other selected times,for example, may be the time for about 25%, 50%, 75%, or 100% (completeerosion) of the device to erode. Complete erosion may correspondapproximately to the time required for treatment by the device.

The “half-life” of a degrading polymer refers to the length of time forthe molecular weight of the polymer to fall to one half of its originalvalue. See e.g., J. C. Middleton and A. J. Tipton, Biomaterials, Vol. 21(23) (2000) pp. 2335-2346.

In addition, metals may be considered to be biostable or bioerodible.Some metals are considered bioerodible since they tend to erode orcorrode relatively rapidly when exposed to bodily fluids. Biostablemetals refer to metals that are not bioerodible. Biostable metals havenegligible erosion or corrosion rates when exposed to bodily fluids.

In general, metal erosion or corrosion involves a chemical reactionbetween a metal surface and its environment. Erosion or corrosion in awet environment, such as a vascular environment, results in removal ofmetal atoms from the metal surface. The metal atoms at the surface loseelectrons and become actively charged ions that leave the metal to formsalts in solution.

Representative examples of bioerodible metals that may be used tofabricate an implantable medical device may include, but are not limitedto, magnesium, zinc, and iron. In one embodiment, a bioerodible metallicstent may be completely eroded when exposed to bodily fluids, such asblood, between about a week and about three months, or more narrowly,between about one month and about two months.

As indicated above, implantable medical devices, such as a stent, shouldbe capable of exhibiting relatively high strength and rigidity, as wellas flexibility since devices have varied mechanical requirements duringuse arising from stress imposed on the device, both before and duringtreatment. “Use” includes manufacturing, assembling (e.g., crimping astent on balloon), delivery of a stent through a bodily lumen to atreatment site, and deployment of a stent at a treatment site. Forexample, a stent requires radial or hoop strength and rigidity to resistradial compressive forces.

The stress imposed on a stent, for example, during use subjectsindividual structural elements to stress. During deployment, thescaffolding and/or coating of a stent can be exposed to stress caused bythe radial expansion of the stent body. In addition, the scaffoldingand/or coating may be exposed to stress when it is mounted on a catheterfrom crimping or compression of the stent. After deployment, radialcompressive forces subject scaffolding and/or coating to stress. Thesestresses can cause the scaffolding to fracture. Failure of themechanical integrity of the stent while the stent is localized in apatient can lead to serious risks for a patient. For example, there is arisk of embolization caused by a piece of the polymeric scaffoldingand/or coating breaking off from the stent.

Conventional methods of constructing a stent from a polymer materialinvolve forming a polymer tube based on a single polymer or polymerblend and then laser cutting a pattern into the tube. Alternatively, apolymer tube may be formed from sheets or films that are rolled andbonded. Polymer tubes and sheets may be formed by various methods,including, but not limited to extrusion, injection molding, or blowmolding.

In extrusion, a polymer melt of a single polymer or polymer blend isconveyed through an extruder which is then formed into a tube. Extrusiontends to impart large forces on the molecules in the axial direction ofthe tube due to shear forces on the polymer melt. The shear forces arisefrom forcing the polymer melt through a die and pulling and forming thepolymer melt into the small dimensions of a tube. As a result, polymertubes formed by conventional extrusion methods tend to possess asignificant degree of axial polymer chain alignment. However, suchconventionally extruded tubes tend to possess no or substantially nopolymer chain alignment in the circumferential direction.

Due to stresses imposed on an implantable medical device during use, itis important for the mechanical stability of a device to have anadequate magnitude of strength both in axial and circumferentialdirections. The direction of stress in structural members can be invarious directions between axial and circumferential. Therefore, anadequate balance of axial and circumferential strength is also importantfor mechanical stability. The relative amount of axial andcircumferential orientation may depend on a number of factors such asthe stent pattern.

A radially expandable device, such as a stent, without an adequatemagnitude and balance of strength in axial and circumferentialdirections may tend to be more prone to mechanical instability. Forexample, a stent made from a tube with an adequate magnitude and balanceof strength in the radial and axial directions may be less susceptibleto cracking during crimping and deployment. Therefore, it may bedesirable to fabricate an implantable medical device with desiredstrength and balance in the axial and circumferential directions.

Desired strength in both directions may be achieved in a number of ways.Some embodiments may include fabricating a device from a tube of acomposite structure with fibers mixed with a continuous phase. In otherembodiments, polymer chain alignment may be induced along thecircumferential direction to increase strength.

In certain embodiments, a polymer for a fiber may be selected that canform crystalline regions with a high modulus and the polymer for acontinuous phase may include a relatively flexible polymer.Representative examples of polymers that may be used for fiberreinforcement include, but are not limited to, poly(L-lactide) andpolyglycolide. Representative polymers that may be used for a continuousphase may include, but are not limited to, poly(DL-lactide) andpoly(ε-caprolactone).

An implantable medical device, such as a stent, with an adequatemagnitude and balance of both circumferential strength and modulus maybe less susceptible to cracking during the crimping process. Inaddition, increased circumferential strength and modulus may allow adecrease in strut width, or generally, a decrease in form factor of astent. Implantable medical devices fabricated from tubes with adequatestrength in both the axial and circumferential directions may possessmechanical properties similar to or better than metal stents with anacceptable wall thickness and strut width.

It is well know by those skilled in the art that molecular orientationor alignment of polymer chains in a polymer is a particularly importantphenomenon that strongly influences bulk polymer properties. Forexample, strength, modulus, yield stress behavior, and elongation tobreak are a few of the important properties that may be influenced byorientation of polymer chains in a polymer. Orientation refers to thedegree of alignment of polymer chains along a longitudinal or covalentaxis of the polymer chains. The degree of molecular orientation in apolymeric material may be induced by applying stress along a preferreddirection.

Polymers in the solid state may have amorphous regions and crystallineregions. Crystalline regions include highly oriented polymer chains inan ordered structure. An oriented crystalline structure tends to havehigh strength and high modulus (low elongation with applied stress)along an axis of alignment of polymer chains. On the other hand,amorphous polymer regions include relatively disordered polymer chainsthat may or may not be oriented in a particular direction. However, ahigh degree of molecular orientation may be induced even in an amorphousregion. An oriented amorphous region also tends to have high strengthand high modulus along an axis of alignment of polymer chains.Additionally, for some polymers under some conditions, inducedorientation in an amorphous polymer may be accompanied bycrystallization of the amorphous polymer.

A polymer may be completely amorphous, partially crystalline, or almostcompletely crystalline. A partially crystalline polymer includescrystalline regions separated by amorphous regions. The polymer chainsof the crystalline regions are not all necessarily oriented in the samedirection. However, a high degree of orientation of crystallites may beinduced by applying stress to a semi-crystalline polymer. The stress mayalso induce orientation of polymer chains in amorphous regions of apolymer.

Polymer tubes fabricated in a conventional manner using extrusion, blowmolding, or injection molding based on a single polymer or polymer blendtend to have inadequate strength and rigidity in the circumferentialdirection. This is due to low polymer chain alignment in thecircumferential direction.

Various embodiments of the present invention include implantable medicaldevices, such as stents, and methods of fabricating such devices from acomposite including a continuous phase and a discrete phase. Thecontinuous phase may include a polymeric matrix and the discrete phasemay include fibers mixed, dispersed, and/or embedded in the matrix.Additionally, either or both the continuous phase or the discrete phasemay include an active agent.

In some embodiments, the discrete phase may include radiopaquematerials. The radiopaque materials may include, for example, metals;alloys; or mixtures of polymers and metal or alloys. In one embodiment,the discrete phase may include metallic fibers, wires, bands, or strips.The metals may include erodible metals, biostable metals, or mixtures ofbiostable and bioerodible metals. Representative metals that may be usedin the discrete phase may include, but are not limited to, magnesium,zinc, iron, platinum, and gold.

A “fiber” may be defined as a unit of matter having a lengthsubstantially longer than its width or diameter. As used herein, a fibercan include, but is not limited to, a filament, a strip, or a wire.

In some embodiments, a polymeric fiber may be formed using any of anumber of methods known in the art including, but not limited to, meltspinning, wet spinning, dry spinning, gel spinning, electrospinning, oran atomizing process. Fibers may be fabricated with relatively highpolymer chain orientation along the fiber axis, and thus relatively highstrength and stiffness.

“Spinning” of polymeric fibers generally involves the extrusion orforcing of a thick, viscous fluid, which is either a polymer melt orsolution, through the tiny holes of a device called a spinneret to formcontinuous filaments of semi-solid polymer. The spinneret has amultiplicity of holds through which polymer melt or solution passthrough. In their initial state, the fiber-forming polymers are solidsand therefore must be first converted into a fluid state for extrusion.This is usually achieved by melting, if the polymers are thermoplastic(i.e., they soften and melt when heated), or by dissolving them in asuitable solvent if they are non-thermoplastic. If they cannot bedissolved or melted directly, they must be chemically treated to formsoluble or thermoplastic derivatives.

In melt spinning, the fiber-forming polymer is melted for extrusionthrough the spinneret and then solidified by cooling. Wet spinninginvolves forming a fiber from a polymer dissolved in a solvent. Thepolymer solution is pumped through a spinneret that is submerged in achemical bath. The dissolved polymer is immiscible in the chemical bath.As the filaments emerge from the spinneret, the polymer precipitatesfrom solution and solidifies.

Dry spinning also involves forming fibers from a polymer solution. Thepolymer solution is pumped through the spinneret. However, instead ofprecipitating the polymer by dilution or chemical reaction,solidification is achieved by evaporating the solvent in a stream of airor inert gas.

Gel Spinning is a type of wet spinning, but is a special process used toobtain high strength or other special fiber properties. In this process,ultra-high molecular weight polymer is dissolved in a solvent at verylow concentration. The concentration is much lower than that typicallyused in wet spinning and dry spinning processes. The polymers or fibersprecipitate from solution and solidify in a chemical bath or in achilled water bath. The fiber is then drawn to orient the polymermolecules. The draw-down ratio is also typically much higher than forwet spinning and dry spinning processes.

The draw-down ratio is defined as the ratio of the length of a drawnfiber to the original length of the fiber. The draw-down ratio for gelspinning can be up to 40:1, while the drawn-down ratio for wet or meltspinning can be about 3-15:1.

The draw-down ratio is defined as the ratio of the length of a drawnfiber to the length of an as-spun fiber. The as-spun fiber refers to asolidified fiber formed from solution or melt. The draw-down ratio forgel spinning can be up to 40:1, while the drawn-down ratio for wet ormelt spinning can be about 3-15:1.

In a dry-jet-wet spinning method, the polymer is not in a true liquidstate during extrusion. The polymer chains are bound together at variouspoints in liquid crystal form. The chains are not completely separated,as they would be in a true solution. This produces strong inter-chainforces in the resulting filaments that can significantly increase thetensile strength of the fibers. In addition, the liquid crystals arealigned along the fiber axis by the shear forces during extrusion. Thefilaments emerge with a relatively high degree of orientation relativeto each other, further enhancing strength. The filaments first passthrough air and then are cooled further in a liquid bath. The draw-downratio in dry-jet-wet spinning is typically less than 1.03:1.

Electrospinning and atomizing processes may be used to producenanofibers. A “nanofiber” refers to a fiber with a dimension in therange of about 1 nm to about 10,000 nm. Electrospinning makes use ofelectrostatic and mechanical force to spin fibers from the tip of a fineorifice or spinneret. In electrospinning, a polymer is dissolved in asolvent or a polymer melt and is placed in a spinneret (e.g., a glasspipet) sealed at one end. The spinneret is maintained at positive ornegative charge by a power supply, for example. When the electrostaticrepelling force overcomes the surface tension force of the polymersolution or melt, the liquid spills out of the spinneret and forms anextremely fine continuous filament.

The strength and modulus of the spun fibers may be increased by drawing.Drawing involves applying tension along the fiber axis. Fibers may bedrawn while extruded fibers are solidifying and/or after they havehardened. Drawing tends to pull the molecular chains together and orientthem along the fiber axis, creating a considerably stronger and rigidfiber along the fiber axis.

As indicated above, favorable mechanical and degradation properties of astent may be obtained by fabricating the stent as a composite.Individual characteristics of the stent (i.e., rigidity, strength,longitudinal flexibility, degradation rate) may be provided by one ormore of the macroscopically distinct materials that make up thecomposite. Thus, one benefit of a composite structure is that individualcharacteristics of a stent may be tuned independently or moreindependently than a stent fabricated from a single polymer or blend.

To provide strength and rigidity to a stent, the fiber may be fabricatedto be relatively strong and stiff with a high modulus along the fiberaxis. The continuous phase may be configured to have differentproperties than the discrete fiber phase. For instance, the continuousphase may be configured to have a lower modulus, and thus greaterflexibility than the discrete phase. Therefore, the continuous phase maybe configured to provide the required flexibility for the stent. Asindicated above, polymers below their T_(g) tend to be relativelybrittle or inelastic and are more flexible and more easily deformed thanabove their T_(g). Therefore, a flexible continuous phase may beobtained by using polymers with a T_(g) above a body temperature.

Additionally, the degradation behavior of the fiber and continuousphases may be configured to have various combinations. In oneembodiment, degradation rates of the fiber and continuous phase may beapproximately the same. Alternatively, the degradation rates of thefiber and the continuous phases may be different. The degradation rateof the fiber may be faster or slower than the continuous phase. Asdiscussed herein, the degradation rate of phases may be controlled thechoice of polymer, the molecular weight, and the crystallinity of thepolymers of the phases.

Moreover, there are numerous ways that the properties of polymers in thestent may be controlled or modified. These include a suitable choice ofpolymers or chemical component groups in polymers for the discrete andcontinuous phases since different polymers have different mechanicalproperties and degradation rates. In addition, as described below,various properties depend on the molecular weight of polymers. Inaddition, certain properties of polymers are also related to the degreeof crystallinity in a polymer. Thus, these properties of the discreteand continuous phases may be modified independently by choice ofpolymers and modifying the molecular weight and crystallinity of thepolymers in these phases.

Mechanical properties such as strength and modulus, the degradationbehavior of polymer, and melting temperatures depend upon the molecularweight. In general, the higher the molecular weight, the stronger andstiffer (higher modulus) a polymer is. Therefore, the strength andmodulus of a fiber may be further enhanced by fabricating a fiber with ahigher molecular weight.

Additionally, the degradation rate of a polymer decreases as themolecular weight increases. Also, the melting temperature increases withmolecular weight. In some embodiments, the same type of polymer withdifferent molecular weights may be used for both the fiber andcontinuous phase in a composite for a device. Due to the differentmolecular weight of the continuous and discrete phases, the mechanicalproperties, degradation behavior, and melting temperature of the phasesmay be different.

As indicated above, the degree of crystallinity in a polymer is relatedto the mechanical properties such as the strength and modulus of amaterial. The higher the degree of crystallinity, the stronger andstiffer a polymer is along the direction of molecular orientation ofcrystalline structures in the polymer.

In addition, the degree of crystallinity is also related to thediffusion rate of fluids, and hence, the erosion rate of a biodegradablepolymer. In general, the diffusion rate of a fluid through a polymerdecreases as the degree of crystallinity increases. Therefore, it isexpected that the diffusion rate of water and bodily fluids is lower incrystalline and semi-crystalline polymers than in amorphous polymers.Thus, the erosion rate of a biodegradable polymeric region may becontrolled by modifying the degree of crystallinity in a continuouspolymeric phase of a composite, for example.

In one embodiment, the crystallinity of a polymer may be modified byheating the polymer. Heating a polymer can alter the degree ofcrystallinity and/or size of crystalline regions in a polymer material.The degree of crystallinity may be altered by heating the polymer withina particular temperature range. Heating a polymer material to atemperature below the glass transition temperature, T_(g), of thepolymer does not significantly alter the molecular structure, and hence,the mechanical properties of the material. Below T_(g), energy barriersto segmental motion of the chains of a polymer inhibit or preventalteration of molecular structure of a polymeric material.

In general, crystallization may occur in a polymeric material that isheated to a temperature between T_(g) and the melting temperature,T_(m), of the polymer. As a result, heating a polymer to a temperaturebetween the T_(g) and the T_(m) of the polymer increases the modulus ofthe polymer.

FIG. 2 depicts a schematic plot of the rate of crystallization of apolymer as a function of temperature. (Rodriguez, F., Principles ofPolymer Systems, 2^(nd) ed., McGraw Hill (1982)) FIG. 2 shows that therate of polymer crystallization increases as the temperature isincreased from below the T_(g) of the polymer or is decreased from abovethe T_(m) of the polymer. The rate of crystallization reaches a maximum16 somewhere between the T_(g) and the T_(m). FIG. 2 shows thateffectively no crystallization occurs below the T_(g) or above theT_(m).

In addition, as indicated above, an amorphous polymer may be formed byheating a polymer material. Above the T_(m), a polymeric materialbecomes a disordered melt and cannot crystallize and any crystallinitypresent is destroyed. Quenching a polymer melt from above the T_(m) ofthe polymer to a temperature below the T_(g) of the polymer may resultin the formation of a solid amorphous polymer. The resulting amorphouspolymer material may have a lower modulus and be a more flexible or aless stiff material than before heating.

In certain embodiments, a method of fabricating a fiber-reinforced stentmay include forming a mixture including a matrix polymer and a pluralityof short or staple fibers. The fibers may include a material having amelting temperature greater than a melting temperature of the matrixpolymer.

In one embodiment, the matrix polymer may be a biostable orbiodegradable polymer or a combination thereof. The material of thefibers may also be a biostable or biodegradable polymer or a combinationthereof. In some embodiments, the material of the fibers may be abiostable and/or erodible metal. In an embodiment, the fibers may becombination of a polymeric and a metallic material. For example, thefibers may be a mixture of polymeric and metallic particles.

As described above, the mixture formed may be a composite material. FIG.3 depicts a schematic representation of the mixture. The matrix polymermay correspond to a continuous phase 20 and short fibers 25 maycorrespond to a discrete phase.

In one embodiment, the short fibers may be composed of the same orsimilar polymeric material as the continuous polymeric phase.Alternatively, the short fibers may be a mixture of fibers withdifferent properties. For example, the short fibers may be a mixture offibers having different degradation rates and/or mechanical properties.

In one embodiment, the mixture may be formed by mixing the matrixpolymer and the fibers in a mixing apparatus at a temperature that isgreater than the melting temperature of the matrix polymer and less thanthe melting temperature of the fiber material. Therefore, a polymericmelt continuous phase containing the matrix polymer may be mixed with adiscrete fiber phase which is below the fiber material meltingtemperature.

FIG. 4 depicts a fiber-reinforced tube with short fibers. Tube 30includes a plurality of short fibers 35 embedded in a continuous polymerphase 38. As shown in FIG. 4, fibers 35 are oriented in arbitrarydirections with respect to the axis of the tube. The fibers providemechanical reinforcement axially, circumferentially, and orientationsbetween the two. Thus, fibers enhance the mechanical stability of thetube and a stent formed from the tube.

In addition, circumferential strength can be further enhanced throughradial expansion of the tube. Radial expansion enhances thecircumferential strength of the tube due to induced polymer chainalignment of the continuous phase and induced circumferential alignmentof the short fibers.

Embodiments of the method may further include disposing the mixture in atube or sheet forming apparatus to form a tube or a sheet. The apparatusmay be heated so that a temperature of the mixture in the apparatus isgreater than the melting temperature of the matrix polymer and less thanthe melting temperature of the fiber material. In some embodiments, atleast a portion of the matrix polymer may be a polymer melt. Inaddition, the mixture may then be cooled below the melting temperatureof the matrix polymer.

As indicated above, a polymer melt may be cooled in such a way so as tocontrol the degree of crystallinity of the formed tube. Thus, in someembodiments, the formed tube or sheet may be cooled to a temperaturebelow the melting temperature of the matrix polymer such that a majorityof the matrix polymer in the formed tube is either amorphous orcrystalline.

In one embodiment, the forming apparatus may be an injection moldingapparatus. The mixture may be injected at a temperature above themelting temperature of the matrix polymer and less than the meltingtemperature of the fiber such that at least a portion of the matrixpolymer is a polymer melt. The mold may be heated by a heating device orin a chamber so that the temperature of the mixture in the mold is abovethe melting temperature of the matrix polymer and less than the meltingtemperature of the fibers.

Alternatively, in another embodiment, the mixture may be placed into themold at a temperature below the melting temperature of the matrixpolymer. The heated mold may then melt the matrix polymer by heating themixture to a temperature above the melting temperature of the matrixpolymer and less the melting temperature of the fibers.

In another embodiment, the forming apparatus may be an extruder. Themixture may be conveyed into the extruder at a temperature below themelting temperature of the matrix polymer. Alternatively, the mixturemay be conveyed into the extruder at a temperature above the meltingtemperature of the matrix polymer and less than the melting temperatureof the fiber such that at least a portion of the matrix polymer is apolymer melt. The mixture may be heated in the extruder so that itstemperature is above the melting temperature of the matrix polymer andless than the melting temperature of the fiber.

Additionally, the method may further include fabricating a stent fromthe tube or sheet. As indicated above, a stent may be fabricated from atube by forming a pattern on the tube including a plurality ofinterconnecting structural elements. Also, a sheet may be rolled into atube and a pattern may be formed onto the tube.

FIG. 5 depicts a schematic representation of an embodiment of a methodof fabricating a fiber reinforced tube, as described above. A matrixpolymer 40 is conveyed into an extruder 45 as either a polymer melt or asolid. Extruder 45 melts and mixes polymer 40 to form a relatively lowviscosity fluid 50. Fluid 50 is fed from extruder 45 into a mixingapparatus 55. Fibers 60 are also fed into mixing apparatus 55. Mixingapparatus 55 mixes fluid 50 with fibers 60 to produce mixture 65.Mixture 65 is conveyed through a die 70 into a forming apparatus 75 toform a tube or a sheet. Forming apparatus 75 may be, for example, aninjection molding apparatus or an extruder.

In some embodiments, the short fibers may be made by forming fibers asdescribed above, and cutting them into short lengths. In one embodiment,a length of at least a portion of the short fibers is substantiallysmaller than a diameter of the formed tube. For example, in someembodiments, the short fibers may be less than 0.05 mm long. In otherembodiments, the short fibers may be between 0.05 and 8 mm long or morenarrowly between 0.1 and 0.4 mm long or 0.3 and 0.4 mm long.

In other embodiments, a method of fabricating a fiber-reinforced stentmay include forming a tube including at least one fiber layer and atleast one polymer film layer. In one embodiment, at least one fiberlayer alternates with at least one film layer. In an embodiment, fibersin a fiber layer may include at least one material having a meltingtemperature greater than melting temperatures at least one polymer filmlayer. Alternatively, the method may include forming a layered sheetincluding at least one fiber layer and at least one polymer film layer.

In one embodiment, at least one polymer film layer may include abiodegradable polymer. The material of the fibers of at least one fiberlayer may include a biostable or biodegradable polymer or a combinationthereof. In an embodiment, the material of the fibers of at least onefiber layer may include a biostable and/or erodible metal. In anembodiment, the fibers of at least one fiber layer may be a combinationof a polymeric and a metallic material. For example, the fibers may be amixture of polymeric and metallic particles.

In one embodiment, the fibers of the fiber layers may be composed of thesame or similar polymeric material as the polymer film layers.Alternatively, the fiber layers may be a mixture of fibers withdifferent properties. Some embodiments may include at least one fiberlayer with different properties than another fiber layer. For example,different fiber layers may have different degradation rates and/ormechanical properties.

In other embodiments, the polymer film layers may have the same orsimilar properties. Alternatively, at least one polymer film layer mayhave different properties than another polymer film layer. For example,different polymer film layers may have different degradation ratesand/or mechanical properties.

In one embodiment, the fiber layer may be a woven structure. A wovenstructure may refer to any structure produced from between one andseveral hundred or more fibers that are woven, braided, knitted,helically wound, and/or intertwined in any manner, at angles between 0°and 180° degrees with the cylindrical axis of the tube, depending uponthe overall geometry and dimensions desired.

FIG. 6A depicts a two-dimensional radial cut-off view of a tube formedwith a fiber layer 80 between two polymer film layers 82 and 84. FIG. 6Bdepicts an expanded view of the layers. Fiber layer 80 is at leastpartially embedded in polymer from polymer film layers 82 and 84 due tomelting of the polymer film layers.

FIG. 7 depicts a tube 90 of helically wound fiber mesh including twosets of helically wound fibers 92 and 94. Tube 90 has a cylindrical axis96. Coordinate system 98 shows the relative orientation with respect toaxis 96. Fibers 92 have a relative orientation greater than 90° andfibers 92 have a relative orientation less than 90°.

In some embodiments, an orientation of fibers in one fiber layer may bedifferent from an orientation of fibers in another fiber layer. This mayfurther enhance the mechanical stability of stent. One embodiment mayinclude one fiber layer with a set of fibers with an orientation greaterthan 90° and another fiber layer with a set of fibers with anorientation less than 90°.

For example, FIG. 8 depicts a two-dimensional view of layers of a tubeformed with fiber layers 100 and 104 and polymer film layers 108, 112,and 116. Fiber layer 100 may include fibers with an orientation greaterthan 90°, such as fibers 92 in FIG. 7. Fiber layer 104 may includefibers with an orientation less than 90°, such as fibers 94 in FIG. 7.

In one embodiment, the tube may be formed by disposing the layers over amandrel. For example, FIG. 9 depicts a helically wound fiber mesh 120disposed on a mandrel 124. A polymer film layer may be disposed overmandrel 124 before disposing fiber mesh 120 over mandrel 124. A polymerfilm layer may be disposed over mandrel 124 followed by another fiberlayer, another polymer film layer, and so on.

Additionally, the method may further include heating the tube or sheetto a temperature greater than the melting temperatures of at least onepolymer film layer and less than the melting temperature of a materialof the fibers. Heating the tube or sheet may melt at least a portion ofthe polymer of the polymer film layers.

In one embodiment, at least a portion of the fiber layers may becomeembedded within at least a portion of the melted polymer of a polymerfilm layer. In some embodiments, the heated tube may be cooled and astent may then be fabricated from the cooled tube. As described above,the heated tube may be cooled in such a way to control the degreecrystallinity of the cooled polymer film layers of the formed tube.

As indicated above, nanofibers may be used in fabricating the stent.Nanofibers are particularly desirable when fabricating a layeredstructure since a larger number of layers may be formed. In general, themore the number of layers, the stronger the composite structure. Thenumber of layers may be limited if fibers larger than nanofibers areused since the structure may become thicker than desirable.

In other embodiments, a method of fabricating a fiber-reinforced stentmay include forming a coating layer including a coating polymer over atube-shaped fiber layer having a plurality of fibers. Alternatively, thefibers may be formed into a sheet. The plurality of fibers shaped into atube may be a woven structure, as described above.

In one embodiment, the coating polymer may include a biostable and/orbiodegradable polymer or a combination thereof. A material of the fibersof at least one fiber layer may include a biostable and/or biodegradablepolymer or a combination thereof. In an embodiment, the material of thefibers may include a biostable and/or erodible metal. In anotherembodiment, the fibers may be a combination of a polymeric and ametallic material. For example, the fibers may be a mixture of polymerand metallic particles.

In one embodiment, the coating layer may be formed by applying a fluidincluding the coating polymer dissolved in a solvent. The material ofthe fibers may be insoluble or have a relatively low solubility in thesolvent. In some embodiments, the coating may include an active agent.The fluid may include an active agent dissolved or dispersed in thefluid. In an embodiment, the material of the fiber may have a meltingtemperature greater than a melting temperature of the coating polymer.Additionally, all or a majority of the solvent may be removed from theapplied fluid.

Furthermore, the fluid may be applied on the tube in a variety of waysknown in the art. For example, the fluid may be sprayed on the tube orthe tube may be dipped in the fluid. In one embodiment, the fiber layermay be disposed on a mandrel and then dipped in and/or sprayed with thefluid.

In one embodiment, the fiber layer may be disposed on a mandrel over apolymer layer including the coating polymer or another type of polymerpreviously formed on the mandrel. The previously formed polymer layer onthe mandrel may be formed by dipping and/or spraying, as describedabove.

In some embodiments, after forming the coating, the tube or sheet may beheated to a temperature above the melting temperature of the coatingpolymer and below the melting temperature of the material of the fiber.The tube or sheet may then be cooled to a temperature below the meltingtemperature of the coating polymer such that a majority of the coatingpolymer in the formed tube or sheet is amorphous, crystalline, orpartially crystalline.

In some embodiments, a method of fabricating a fiber-reinforced stentmay include

disposing a plurality of fibers within a mold for forming a structure.The structure may be, for example, a tube or a sheet. The fibers may bedisposed within the mold in a number of ways. One embodiment may includedisposing short fibers, as described above, in a random or substantiallyrandom fashion within the mold. In another embodiment, long fibers maybe wound around a mandrel, disposed in the mold, in a helical or otherfashion. In one embodiment, a woven structure, as described above, maybe disposed in the mold.

Additionally, the method may further include disposing a matrix polymerthat is partially or completely molten into the mold to at leastpartially embed the fibers within the molten polymer. In one embodiment,the fibers may include a material having a melting temperature greaterthan a melting temperature of the matrix polymer. The temperature of themolten polymer and the fibers within the mold may be less than a meltingtemperature of the material of the fiber.

In one embodiment, the matrix polymer may be a biostable orbiodegradable polymer or a combination thereof. The material of thefibers may also be a biostable or biodegradable polymer or a combinationthereof. In some embodiments, the material of the fibers may be abiostable and/or erodible metal. In an embodiment, the fibers may becombination of a polymeric and a metallic material. For example, thefibers may be a mixture of polymer and metallic particles.

The molten polymer may then be cooled to form the structure and a stentmay be fabricated from the cooled structure. As indicated above, thepolymer melt may be cooled in such a way to control the degreecrystallinity of the matrix polymer.

A stent may be formed from a tube by forming a pattern in the tubeincluding a plurality of interconnecting structural elements. Asindicated above, a stent may be fabricated from a sheet by forming atube from the sheet and forming a pattern in the tube including aplurality of interconnecting structural elements.

In one embodiment, a medicated stent may be fabricated by disposing anactive agent into the mold. The active agent may be mixed or dispersedwithin the molten matrix polymer. Alternatively, the active agent may bemixed or dispersed with the fiber. In another embodiment, a coatingincluding an active agent may be applied to the stent.

In further embodiments, a method of fabricating a fiber-reinforced stentmay include

disposing a plurality of fibers in an extruder for forming a structure.The structure may be, for example, a tube or a sheet. As describedabove, the fibers may be disposed within the extruder in a number ofways. One embodiment may include disposing short fibers, as describedabove, in a random or substantially random fashion within the extruder.In another embodiment, long fibers may be wound around a mandrel,disposed in the mold, in a helical or other fashion. In one embodiment,a woven structure, as described above, may be disposed in the extruder.

Additionally, the method may further include conveying a matrix polymerinto the extruder. In one embodiment, the matrix polymer may have amelting temperature less than a melting temperature of a material of thefiber. In addition, the structure may then be formed with the extruderat a temperature greater than the melting temperature of the matrixpolymer and less than the melting temperature of the material of thefiber. In an embodiment, at least some of the fibers may become embeddedwithin matrix polymer.

In one embodiment, the matrix polymer may be a biostable orbiodegradable polymer or a combination thereof. The material of thefibers may also be a biostable or biodegradable polymer or a combinationthereof. In some embodiments, the material of the fibers may be abiostable and/or erodible metal. In an embodiment, the fibers may becombination of a polymeric and a metallic material. For example, thefibers may be a mixture of polymer and metallic particles.

The molten polymer may then be cooled to form the structure and a stentmay be fabricated from the cooled structure. A stent may be fabricatedfrom a tube or a stent as described above. As indicated above, themolten polymer melt may be cooled in such a way to control the degreecrystallinity of the matrix polymer.

In some embodiments, a medicated stent may be fabricated by conveying anactive agent into the extruder. In other embodiments, at least some ofthe fibers may include an active agent.

As indicated above, a pattern including a plurality of interconnectingstructural elements may be formed by laser cutting the pattern. In oneembodiment, the pattern can be cut such that at least a portion of thestructural elements may be aligned or substantially aligned with anorientation of at least some of the fibers. For example, FIG. 10illustrates a pattern of struts or structural elements 130 that arealigned with the orientation of fiber segments 132, 134, 136, and 138.

In further embodiments, the circumferential strength and rigidity of afiber reinforced stent may be enhanced by radial expansion of a tube.Polymer chain alignment in the continuous polymer phase may be inducedalong the circumferential direction to increase strength. In addition,radial expansion may also induce alignment of fibers in the tube,further enhancing circumferential strength.

The degree of polymer chain alignment induced with applied stress maydepend upon the temperature of the polymer. For example, below the T_(g)of a polymer, polymer segments may not have sufficient energy to movepast one another. In general, polymer chain alignment may not be inducedwithout sufficient segmental mobility. Above T_(g), polymer chainalignment may be readily induced with applied stress since rotation ofpolymer chains, and hence segmental mobility, is possible. Between T_(g)and the melting temperature of the polymer, T_(m), rotational barriersexist. However, the barriers are not great enough to substantiallyprevent segmental mobility. As the temperature of a polymer is increasedabove T_(g), the energy barriers to rotation decrease and segmentalmobility of polymer chains tends to increase. Thus, as the temperatureincreases, polymer chain alignment is more easily induced with appliedstress.

Rearrangement of polymer chains may take place when a polymer isstressed in an elastic region and in a plastic region of the polymermaterial. A polymer stressed beyond its elastic limit to a plasticregion generally retains its stressed configuration and correspondinginduced polymer chain alignment when stress is removed. The polymerchains may become oriented in the direction of the applied stress. Thestressed polymer material may have a higher tensile strength and modulusin the direction of the applied stress.

Additionally, heating a polymer may facilitate deformation of a polymerunder stress, and hence, modification of the mechanical properties ofthe polymer. A polymer deformed elastically with stress facilitated withheating may retain induced polymer chain alignment by cooling thepolymer before relaxing to or towards an unstrained state.

In some embodiments, a polymer tube may be deformed at a temperaturebelow the T_(g) of the polymer of the continuous phase. Alternatively,it may be desirable to deform the tube in a temperature range greaterthan or equal to the T_(g) of the continuous phase polymer and less thanor equal to the T_(m) of the polymer. As indicated above, a polymericmaterial deforms much more readily due to segmental motion of polymerchains above T_(g). Deformation induces polymer chain alignment that mayoccur due to the segmental motion of the polymer chains. Therefore,heating the polymer tube prior to or contemporaneously with deformationmay facilitate deformation particularly for polymers with a T_(g) belowan ambient temperature. Heating the tube contemporaneously with thedeformation may be desirable since the deformation may occur at aconstant or nearly constant temperature. Therefore, the induced polymeralignment and material properties may be constant or nearly constant.

In some embodiments, a fiber-reinforced polymer tube may be deformedradially by increasing a pressure in a polymer tube, for example, byconveying a fluid into the tube. Tension and/or torque may also beapplied to the tube. The tube may be positioned in an annular member ormold. The mold may act to control the degree of radial deformation ofthe tube by limiting the deformation of the outside diameter or surfaceof the tube to the inside diameter of the mold. The inside diameter ofthe mold may correspond to a diameter less than or equal to a desireddiameter of the polymer tube.

The polymer tube may also be heated prior to, during, and subsequent tothe deformation. In general, it is desirable for the temperature duringdeformation to be greater than or equal to a glass transitiontemperature of the polymer and less than or equal to a meltingtemperature of the polymer. The polymer tube may be heated by the fluidand/or the mold.

Certain embodiments may include first sealing, blocking, or closing apolymer tube at a distal end. The end may be open in subsequentmanufacturing steps. A fluid, (conventionally an inert gas such as air,nitrogen, oxygen, argon, etc.) may then be conveyed into a proximal endof the polymer tube to increase the pressure in the tube. The pressureof the fluid in the tube may act to deform the tube.

The increased pressure may deform the tube radially and/or axially. Thefluid temperature and pressure may be used to control the degree ofradial deformation by limiting deformation of the inside diameter of thetube as an alternative to or in combination with using the mold. Inaddition, it may be desirable to increase the pressure to less thanabout an ultimate stress of the continuous phase polymer to inhibit orprevent damage to the tube. The continuous phase polymer may be deformedplastically or elastically. As indicated above, a polymer elongatedbeyond its yield point tends to retain its expanded configuration, andhence, tends to retain the induced molecular orientation.

Additionally, the pressure inside the tube and the temperature of thetube may be maintained above ambient levels for a period of time toallow the polymer tube to be heat set. In one embodiment, thetemperature of the deformed tube may be maintained at greater than orequal to the T_(g) of the continuous phase polymer and less than orequal to the T_(m) of the continuous polymer for a selected period totime. The selected period of time may be between about one minute andabout two hours, or more narrowly, between about two minutes and aboutten minutes. “Heat setting” refers to allowing polymer chains toequilibrate to different configurations in response to an elevatedtemperature. In this case, the polymer chains are allowed to adopthighly oriented structure at an elevated temperature. Polymer chainalignment is a time and temperature dependent process, therefore, aperiod of time may be necessary to allow polymer chains to realign at agiven temperature that are stable in a deformed state of a polymericmaterial. Heat setting may also be facilitated by tension.

Further embodiments of the present invention relate to stents composedprimarily or completely of polymeric fibers coiled or braided into amesh tube or stent structure. A braided stent can provide sufficientradial strength, however, the radial profile of such a mesh, fiberstructure can be higher than desirable. In particular, the “net points,”which refer to the points of overlap of fibers, tend to increase theradial profile of a stent. Net points 142 are illustrated in FIG. 9.Embodiments of methods described herein allow fabrication of a fibermesh stent with a sufficiently small profile and sufficiently highradial strength.

Certain embodiments of a method of fabricating a stent may includemaking a tube or stent structure from at least two types of fibers. Inone embodiment, a first fiber may include a first polymer and a secondfiber may include a second polymer.

Other embodiments of a method of fabricating a stent may include makinga tube or stent structure from composite fibers including the firstpolymer and the second polymer. In some embodiments, the fibers of thetube may include an inner core including the first polymer and an outerlayer or covering including the second polymer. In other embodiments,the outer covering may be the first polymer and the inner core may bethe second polymer. In alternative embodiments, the fibers of the tubemay be a mixture or blend of the first polymer and the second polymer.

In certain embodiments, the first polymer may have a softeningtemperature (T_(s)) lower than a softening temperature of the secondpolymer. Also, the first polymer may have a T_(m) and a T_(g) less thanthe T_(m) and T_(g) of the second polymer.

For example, in a stent including two different types of fibers, a firstfiber may be made from poly (l-lactic acid) and a second fiber may bemade from 10:90 poly(l-lactide-co-glycolide) (10% lactide, 90%glycolide). The poly (L-lactic acid) has a melting temperature of 175°C. and poly(l-lactide-co-glycolide) has a melting temperature of 200° C.Additionally, an exemplary composite fiber may be made from 50% poly(l-lactic acid) and 10:90 poly(l-lactide-co-glycolide).

Various embodiments of fabricating a stent having two types of fibersand/or fibers composed of a mixture of the first polymer and the secondpolymer may include heating the tube to a temperature range between asoftening temperature of the first polymer and the melting temperatureof the first polymer. In one embodiment, pressure is applied to the tubeso as to flatten at least some of the fibers of the tube to reduce theradial profile of the tube. A heated fiber made of a first polymer maytend to flatten as pressure is applied, reducing the profile of thetube. Likewise, a heated fiber including the first polymer may also tendto flatten. In particular, the applied pressure may reduce the radialprofile of the tube of at least some of the net points of the fibers.

In some embodiments, the second polymer may be adapted to provide highstrength to the stent structure during heating, flattening of thefibers, and during use of the stent. In an embodiment, the temperaturerange may be below the T_(g) of the second polymer. In this case, thesecond polymer may remain relatively rigid during heating a flatteningof the first fiber. The temperature range may also be below the T_(s) ofthe second polymer. Alternatively, the temperature range may be abovethe T_(s) of the second polymer.

In some embodiments, a cross-section of a fiber may be circular.Alternatively, a cross-section of a fiber may have an oblong shape, forexample, an oval or elliptical shape. Fibers with an oblong-shapedcross-section may allow greater surface coverage of a vessel, inaddition to providing a smaller radial profile to the stent.

In one embodiment, the tube may be disposed over a mandrel duringheating and flattening of the fibers. Pressure may be applied forflattening the fibers by a pressure tube disposed over the tube. In analternate embodiment, the fibers may be heated and flattened in a heatedcrimper. Heat may be applied to the tube by a heated mandrel. The tubemay also be heated by blowing a heated fluid onto the tube such as aninert gas, e.g., argon, air, oxygen, nitrogen, etc. Additionally, thetube may be allowed to heat set on the mandrel.

In an embodiment, the tube may be heated and the pressure applied at ornear a fabricated diameter of the tube. The tube may be allowed to heatset by maintaining the tube in the temperature range for a selectedperiod of time.

In some embodiments, the method may include radial expansion the tubeprior to, during, or subsequent to heating and/or applying pressure toflatten at least some of the fibers. As described above, radialexpansion may induce molecular orientation in the fiber polymer thattends to increase the tensile strength of the fiber. In otherembodiments, the tube may be crimped prior to, during, or subsequent toheating the tube and/or applying pressure to flatten at least some ofthe fibers.

FIG. 11 depicts a radial cross-section of composite fiber 140. Fiber 140has an inner core 142 of a first polymer and an outer layer 144 of asecond polymer. FIG. 12 depicts a radial cross-section of a system 150that can be used for heating and flattening fibers of a fiber stent.Overlapping fibers 152 are between a sliding wedge-type crimper 154 anda stationary mandrel 156. Wedge-type crimper 154 is heated and can applypressure to the fibers of the stent. A region 160 including fibers 152,crimper 154, and mandrel 156 in FIG. 12 is shown in an expanded view inFIGS. 13 and 14. FIG. 13 illustrates fibers 152 prior to applyingpressure to the fiber by inward movement of crimper 154 and FIG. 14shows fibers 152 after applying pressure with crimper 154. FIG. 14 showsthat fibers 152 have been flattened by crimper 154.

As indicated above, there are difficulties associated with manufacturingstents with small radiopaque markers. Various embodiments of stents andmethods of making stents that include metallic films coupled and/orembedded within polymeric stents are disclosed herein. The metallic filmmay be sufficiently radiopaque to allow the stent to be visualizedduring use.

The stents may be formed by coupling and/or embedding metallic film inand/or on a polymeric tube. A stent may be fabricated by forming apattern of interconnecting structural elements in the tube with themetallic film using, for example, a laser. Forming the pattern mayinclude removal of some of the metallic film in or on the polymer inaddition to removal of polymer.

In certain embodiments, the metallic film may include a biostable metal,a bioerodible metal, or a combination of a biostable and bioerodiblemetal. As indicated above, representative examples of bioerodible metalsthat may be used to fabricate an implantable medical device may include,but are not limited to, magnesium, zinc, and iron. Representativeexamples of biostable metals that may be used to fabricate animplantable medical device may include, but are not limited to, gold orplatinum.

In certain embodiments, a stent may include metallic film coupled to aplurality of portions of a surface of the stent. Some embodiments of amethod of making the stent may include coupling a metallic film to atleast a portion of a surface of a polymeric tube. In one embodiment, themetallic film may include a band circumferentially aligned around asurface of the tube. A length of a band along a longitudinal axis of thestent may be less than or equal to the length of the tube. In anotherembodiment, the metallic film includes a longitudinal striplongitudinally aligned along the surface of the tube.

As an illustration, FIG. 15 depicts a polymeric tube 170 with acircumferentially aligned metallic band 174 coupled or adhered to thesurface of tube 170. In addition, FIG. 16 depicts a polymeric tube 178with a longitudinally aligned strip of metallic film 182 coupled oradhered to the surface of tube 178.

In an alternate embodiment, a method of making a stent may includecoupling the metallic film to at least a portion of a surface of apolymeric sheet. The sheet may then be rolled and bonded to form a tube.

FIG. 17 depicts a portion of a stent 200 fabricated from tube 170 inFIG. 15. Stent 200 has circumferentially aligned metallic film markers204 coupled to structural elements 208. Line A-A corresponds to thelongitudinal axis of the stent. FIG. 18 depicts a stent 220 fabricatedfrom tube 178 in FIG. 16. Stent 220 has longitudinally aligned metallicmarkers 224 coupled to structural elements 228. Line A-A corresponds tothe longitudinal axis of the stent.

In some embodiments, the metallic film may be coupled to the polymerictube using any suitable biocompatible adhesive. In one embodiment, theadhesive may include a solvent. The solvent may dissolve the polymer ofthe polymeric tube to allow the metal film to be coupled to the tube. Inanother embodiment, the adhesive may include a solvent mixed with apolymer. The solvent or the solvent-polymer mixture may be applied tothe tube followed by application of the metallic film. The solvent maythen be removed by heating the tube, for example, in an oven.

Representative examples of solvents may include, but are not limited to,chloroform, acetone, chlorobenzene, ethyl acetate, 1,4-dioxane, ethylenedichloride, 2-ethyhexanol, and combinations thereof. Representativepolymers may include biostable and biodegradable polymers disclosedherein that may be dissolved by the selected solvent.

In other embodiments, adhesives may include, but are not limited to,thermosets such as, for example, epoxies, polyesters and phenolics;thermoplastics such as, for example, polyamides, polyesters and ethylvinyl acetate (EVA) copolymers; and elastomers such as, for example,natural rubber, styrene-isoprene-styrene block copolymers, andpolyisobutylene. Other adhesives include, but are not limited to,proteins; cellulose; starch; poly(ethylene glycol); fibrin glue; andderivatives and combinations thereof.

Mixtures of solvents and another substance can be used to formadhesives. In some embodiments, mixtures of water and sugar such as, forexample, mixtures of water and sucrose, can be used as an adhesive. Inother embodiments, mixtures of PEG, or derivatives thereof, can be mixedwith a suitable solvent to form an adhesive. Suitable solvents for PEG,or derivatives thereof, include, but are not limited to, water, ethanol,chloroform, acetone, and the like.

In other embodiments, the method may further include forming a coatingabove an outer surface of the stent with a metallic film coupled to asurface of the stent. The coating may be above at least a portion of themetallic film on the surface of the stent. In one embodiment, thecoating may include a biostable or biodegradable polymer. In oneembodiment, the coating may include an active agent or drug. The coatingmay be formed by applying a mixture of a polymer and a solvent, followedby removal of the solvent. In an embodiment, the polymer coating mayinhibit or prevent detachment of the metallic film from the stent priorto substantial or complete biodegradation of a biodegradable coating.

As an illustration, FIG. 19 depicts a cross-sectional view of a sidewallof a portion 230 of a structural element of a stent. A coating 232 isabove a metallic film marker 234 which is coupled or adhered to apolymeric substrate 238. Coating 232 tends to inhibit detachment ofmarker 234 from substrate 238.

In other embodiments, structural elements of a stent may include tworadial polymeric layers with metallic film embedded in a plurality oflocations in between the layers. In certain embodiments, a method ofmaking the stent may include forming a tube including a metallic filmembedded between two polymer layers and fabricating a stent from thetube. In an embodiment, the metallic film may be a bandcircumferentially aligned around the tube in between the polymericlayers. In another embodiment, the metallic film may be a longitudinalstrip longitudinally aligned along the tube in between the polymericlayers.

Alternatively, a sheet may be formed including the metallic filmembedded between two polymer layers. The sheet may then be rolled andbonded to form a tube.

As an illustration, FIG. 20 depicts a cross-sectional view of a sidewallof a portion 240 of a structural element of a stent. A metallic filmmarker 244 is embedded between a luminal polymeric layer 248 and anabluminal polymeric layer 252.

Some embodiments may include forming the tube by extruding an outerpolymeric tubular layer over an inner tubular polymeric layer with ametallic film disposed above a surface of the inner layer. Extruding theouter layer over the inner layer may then embed the metallic filmbetween the layers. In some embodiments, the melting temperature of aninner polymeric layer may be higher than the melting temperature of theouter layer. The outer layer may be extruded over the inner layer at atemperature above the melting temperature of the outer polymer layer andbelow the melting temperature of the inner polymer layer, allowing theinner layer to maintain its structural integrity.

In further embodiments, a method of making a stent may includeelongating a polymeric tube so that a diameter of the stent decreases. Ametallic film in the form of a metallic band may then be positionedaround at least a portion of the elongated tube. The polymeric tube withthe metallic band positioned around the tube may then be heated. In someembodiments, the method may further include allowing the heated tube toradially expand so as to couple the metallic band to the tube. Theheated tube may radially expand to at least a diameter of the metallicband.

Representative examples of polymers that may be used to fabricateembodiments of implantable medical devices disclosed herein include, butare not limited to, poly(N-acetylglucosamine) (Chitin), Chitosan,poly(3-hydroxyvalerate), poly(lactide-co-glycolide),poly(3-hydroxybutyrate), poly(4-hydroxybutyrate),poly(3-hydroxybutyrate-co-3-hydroxyvalerate), polyorthoester,polyanhydride, poly(glycolic acid), poly(glycolide), poly(L-lacticacid), poly(L-lactide), poly(D,L-lactic acid), poly(D,L-lactide),poly(L-lactide-co-D,L-lactide), poly(caprolactone),poly(L-lactide-co-caprolactone), poly(D,L-lactide-co-caprolactone),poly(glycolide-co-caprolactone), poly(trimethylene carbonate), polyesteramide, poly(glycolic acid-co-trimethylene carbonate),co-poly(ether-esters) (e.g. PEO/PLA), polyphosphazenes, biomolecules(such as fibrin, fibrinogen, cellulose, starch, collagen and hyaluronicacid), polyurethanes, silicones, polyesters, polyolefins,polyisobutylene and ethylene-alphaolefin copolymers, acrylic polymersand copolymers other than polyacrylates, vinyl halide polymers andcopolymers (such as polyvinyl chloride), polyvinyl ethers (such aspolyvinyl methyl ether), polyvinylidene halides (such as polyvinylidenechloride), polyacrylonitrile, polyvinyl ketones, polyvinyl aromatics(such as polystyrene), polyvinyl esters (such as polyvinyl acetate),acrylonitrile-styrene copolymers, ABS resins, polyamides (such as Nylon66 and polycaprolactam), polycarbonates, polyoxymethylenes, polyimides,polyethers, polyurethanes, rayon, rayon-triacetate, cellulose, celluloseacetate, cellulose butyrate, cellulose acetate butyrate, cellophane,cellulose nitrate, cellulose propionate, cellulose ethers, andcarboxymethyl cellulose. Additional representative examples of polymersthat may be especially well suited for use in fabricating embodiments ofimplantable medical devices disclosed herein include ethylene vinylalcohol copolymer (commonly known by the generic name EVOH or by thetrade name EVAL™), poly(butyl methacrylate), poly(vinylidenefluoride-co-hexafluoropropene) (e.g., SOLEF® 21508, available fromSolvay Solexis PVDF, Thorofare, N.J.), polyvinylidene fluoride(otherwise known as KYNAR®, available from Atofina Chemicals,Philadelphia, Pa.), ethylene-vinyl acetate copolymers, poly(vinylacetate), styrene-isobutylene-styrene triblock copolymers, andpolyethylene glycol.

While particular embodiments of the present invention have been shownand described, it will be obvious to those skilled in the art thatchanges and modifications can be made without departing from thisinvention in its broader aspects. Therefore, the appended claims are toencompass within their scope all such changes and modifications as fallwithin the true spirit and scope of this invention.

What is claimed is:
 1. A method of making a stent comprising: providinga tube comprising a woven structure including composite fibers, whereinthe composite fibers comprise a first polymer and a second polymer,wherein the first polymer and the second polymer are bioabsorbable,wherein the first polymer has a Vicat softening temperature (Ts) lowerthan a Ts of the second polymer; heating the tube to a temperature rangebetween the Ts of the first polymer and the Ts of the second polymer,wherein the temperature range is below a melting temperature of thefirst polymer and the second polymer; and applying pressure to the tubeso as to flatten at least some of the fibers of the tube to reduce aradial profile of the tube.
 2. The method of claim 1, wherein the tubeis disposed over a mandrel during heating.
 3. The method of claim 1,wherein pressure is applied with a heated crimper.
 4. The method ofclaim 1, wherein a diameter of the tube is fixed during heating.
 5. Themethod of claim 1, wherein the tube is heated and pressure applied at ornear a fabricated diameter of the tube.
 6. The method of claim 1,wherein the temperature of the tube is maintained in the temperaturerange for a selected period of time to allow heat setting of the tube.7. The method of claim 1, further comprising radially expanding the tubeprior to, during, or subsequent to heating and/or applying pressure toflatten at least some of the fibers.
 8. The method of claim 1, whereinthe woven structure comprises fibers that are braided, knitted,helically wound, and/or intertwined.
 9. The method of claim 1, whereinthe applied pressure reduces a radial profile of at least some netpoints of the fibers.
 10. The method of claim 1, wherein the compositefibers have an inner core of the first polymer and an outer layer of thesecond polymer.